High energy ionizing radiation such as gamma radiation cannot be detected directly; it must be converted into an electrical pulse. One common method for creating an electrical pulse when ionizing radiation is present is to first absorb the ionizing radiation in a scintillator. In response, the scintillator then produces a flash of light which is converted into an electrical signal by a photodetector.
Although the physical configuration of the scintillator/ photodetector combination varies from application to application, the underlying principle remains constant. The scintillator will vary in size, but must be thick enough to stop the incident high energy radiation and large enough to cover the desired area. The exact type of radiation detector that the scintillator is coupled to varies but the radiation detector must produce an electrical signal large enough to be observable above background noise.
The effectiveness of this method of high energy radiation detection is primarily limited by the scintillating material. Only a few materials are known to scintillate, and the scintillation properties of these materials vary. An ideal scintillator would have a high density, a short decay time constant (i.e. the photons are emitted as soon as possible after the radiation interacts in the scintillator), and a large light output that is essentially proportional to the amount of high energy radiation deposited in the scintillator. High density is desirable in order to stop the ionizing radiation in as short a distance as possible, a short decay time is desirable in order to measure the time of interaction accurately, and a high light output is desirable in order to make it easier for the radiation detector to convert the light into an electrical pulse whose size indicates the amount of energy of radiation. In addition, the scintillating materials should not have unpleasant chemical or material properties, such as toxicity, hygroscopy, or extreme reactivity. Moreover, it is advantageous to utilize a material that is readily fabricated in crystal form.
The radiation detector that converts the scintillation light into an electrical pulse is usually either a photomultiplier tube or a solid-state (Si, GaAs, HgI.sub.2, etc.) photodiode. A photomultiplier tube (PMT) is usually employed when very small amounts of radiation are to be detected, as a PMT converts each light into electrons with a gain of about 1 million. Photodiodes are usually used when large amounts of radiation are to be detected, as a photodiode typically converts light into electrons with unity gain. These radiation detectors can either be coupled to the scintillator either singly or in position sensitive arrays.
FIG. 1 diagrammatically demonstrates a single embodiment of a high energy radiation detector contemplated in the present invention. The device 10 is in the form of a hand held radiation monitor and is similar to a standard Geiger counter, but will be more sensitive to gamma radiation. A scintillator crystal 11 is optically coupled to a photomultiplier tube 12 and this assembly is encased in a opaque material 13 in order to shield ambient light. A cable 14 connects the scintillator/radiation detector assembly to a small box 15 containing a battery operated power supply for the photomultiplier tube, counting electronics, and a numerical display. When ionizing radiation 17 from a source 16 impinges on the scintillator crystal, it emits light which is converted into an electrical pulse by the photomultiplier tube 12 and the rate at which these pulses arrive is determined by the electronic circuit in the box 15 and displayed accordingly. As this rate is proportional to the amount of radiation present, the device can be used to monitor radiation.
A more complicated embodiment of a radiation detector is a PET camera. PET (positron emission tomography) is a medical imaging technique in which a radioactively labeled substance is administered to a patient and then traced within the patient's body by means of an instrument that detects the decay of the isotope. In PET, a chemical tracer compound having a desired biological activity or affinity for a particular organ is labeled with a radioactive isotope that decays by emitting a positron (positive electron). The emitted positron stops after traveling only a few millimeters in living tissue. It then interacts with an electron, an event that annihilates both particles. The mass of the two particles is converted into 1.02 million electron volts (1.02 MeV) of energy, divided equally between two 511 keV photons (gamma rays). The two photons are emitted simultaneously and travel in opposite directions. The two photons penetrate the surrounding tissue, exit the patient's body, and are absorbed and recorded by radiation detectors typically arranged in a number of circular arrays.
Biological activity within an organ under investigation can be assessed by tracing the source of the radiation emitted from the patient's body to the radiation detectors. The source of the radiation can be accurately estimated by linking each radiation detector with several other radiation detectors on the opposite side of the radiation detector array and registering a signal only if two detectors sense 511 keV photons within typically 10 nsec. When a coincidence is registered, an annihilation is recorded along a line connecting the two radiation detectors. In this manner, a circumferential array of radiation detectors can establish the sources of all coincident pairs of photons that originate within a volume defined by straight lines joining paired detectors. A computer program reconstructs the spatial distribution of the decaying isotopes within the patient. With suitable interpretation, PET images provide a noninvasive, regional assessment of many biochemical processes associated with human organs.
The value of PET as a clinical imaging technique is in large measure dependent upon the performance of the radiation detectors. The typical PET camera comprises an array of radiation detectors consisting of scintillator crystals coupled to photomultiplier tubes (PMT's). When a high energy photon strikes a detector, it produces light in one of the scintillator crystals that is then sensed by the PMT, which registers the event by passing an electronic signal to the reconstruction processing circuitry.
As pointed out earlier, the scintillator crystals used in the PMT of a general purpose detector or a PET camera must have certain properties, among which are (1) good stopping power, (2) high light yield, and (3) fast decay time.
In a PET application, stopping power is the ability to stop the 511 keV photons in as little material as possible so as to reduce the overall size of the radiation detector, which reduces the cost and improves spatial accuracy. In other words, good stopping power allows for the use of more radiation detectors in a given space, and a corresponding increase in resolution. Stopping power is typically expressed as the linear attenuation coefficient (tau) having units of inverse cm.sup.-1. After a photon beam has traveled a distance "d" in a crystal, the fraction of photons that have not been stopped by the crystal is calculated as follows: EQU fraction of unstopped photons=e.sup.-dtau.
Therefore, after traveling a distance of 1/tau (the "absorption length"), approximately 37% of the photons will not have been stopped and 63% will have been stopped. Likewise, 63% of the remaining photons will have been stopped after traveling an additional distance of 1/tau. For PET, one wants 1/tau to be as small as possible so that the radiation detector is as compact as possible.
Light yield is also an important property of scintillators contemplated for use in PET. Light yield is sometimes referred to as light output or relative scintillation output, and is typically expressed as the percentage of light output from a `standard` crystal such as thallium-doped sodium iodide. Accordingly, the light yield for NaI(T1) is defined as 100.
A third important property of scintillators in PET applications is decay time. Scintillation decay time, sometimes referred to as the time constant or decay constant, is a measure of the duration of the light pulse emitted by a scintillator, and is typically expressed in units of nanoseconds (nsec). As noted above, in PET, the source of biological activity within an organ under investigation is determined by tracing the source of coincident photons emitted from the patient's body to the radiation detectors. When two 511 keV photons are detected at the same time by a pair of radiation detectors, the source of the photons is known to lie along the linear path connecting the two radiation detectors. In general, only a fraction of the detected photons are in coincidence and thus used in the reconstruction analysis. Moreover, many false coincidences are registered because the finite decay time associated with each scintillator may cause it to emit light at the same time as another scintillator when in fact the photons inducing the light did not come from the same positron annihilation. For example, a photon arrived at one radiation detector may produce a flash of light that does not decay, i.e. "turn off", until after a later photon, from a different positron annihilation, produces a flash of light in a detector on the side opposite the first detector. In this instance, the flashes would overlap, the radiation detectors would register them as in coincidence. Thus, scintillator materials with long decay constants have an inherent problem in detecting coincident photons. This also has an adverse effect on the resolution.
For medical reasons, the positron emitting tracer compounds should have very short half-lives. Because of the short halflives of these compounds, data on the occurrence of coincident photons needs to be gathered at as high a rate as possible. As noted above, a long decay time will lead to false coincidences, and therefore each radiation detector must be turned off for a length of time corresponding to its decay time. Since the detector is inactive for this time, it is not collecting data, and reduction of such time increases the amount of data that may be collected and thereby increases the resolution.
In addition to the three important properties discussed above, scintillator crystals for PET should be easy to handle. For example, certain known scintillators are very hygroscopic, i.e., they react with moisture, making it necessary to very tightly encapsulate them to allow their use as scintillators in PET. These hygroscopic scintillators are expensive and difficult to use.
Known scintillator materials include (1) plastic scintillators, (2) thallium-doped sodium iodide (NaI(T1)), (3) cesium fluoride (CsF), (4) bismuth germanate (Bi.sub.4 Ge.sub.3 O.sub.12, also referred to as "BGO"), (5) cerium fluoride (CeF.sub.3), and (6) barium fluoride (BaF.sub.2). Of these six scintillators, only two, BGO and BaF.sub.2, are used routinely for PET.
Plastic scintillators, typically composed of polystyrene doped with a wavelength-shifting additive, are commercially available under such tradenames as PILOT U and NE 111. Upon excitation with a 511 keV photon, plastic scintillators emit a light pulse having a very fast decay constant of approximately 1.5 nsec and light output proportional to the energy of the incident photon. The main disadvantage of plastic scintillators is their low density (approximately 1.1 to 1.2 g/cm.sup.3) due to the light atoms (hydrogen and carbon) that make up the molecules of the material. Because of their low density, plastic scintillators have poor stopping power, and are therefore poorly suited for use in PET.
NaI(T1), thallium-doped sodium iodide, has the highest light output of the six scintillators listed above. NaI(T1) also has reasonably good stopping power (1/tau=3.0 cm at 511 keV). However, NaI(T1) has a long decay constant (250 nsec), a significant disadvantage for use in PET. NaI(T1) has an additional disadvantage in that it is highly hygroscopic, making it extremely difficult to handle in that it must be tightly encapsulated in bulky cans.
CsF, cesium fluoride, has an advantage over plastic scintillators because of its relatively high density (4.61 g/cm.sup.3) and consequent stopping power. However, the light output and decay constant of CsF are inferior to those of plastic scintillators. CsF is also highly hygroscopic, well above NaI(T1) which, as noted above, makes it expensive and difficult to handle.
BGO has the highest density (7.13 g/cm) of the six known scintillator materials noted above. Its stopping power is the best of the six materials (1/tau=1.1 cm at 511 keV). As a result, BGO is best able to absorb 511 keV photons efficiently in small crystals. However, BGO's very long decay constant (300 nsec), longer even than NaI(T1), is at a significant disadvantage for use in PET.
The use of BaF.sub.2 as a scintillator material is described in Allemand et al. U.S. Pat. No. 4,510,394. BaF.sub.2 emits light having two components: 75% is emitted with a `slow` decay constant of approximately 620 nsec and 25% with a `fast` decay constant of approximately 0.6 nsec. BaF.sub.2 has a light yield of approximately 16% that of NaI(T1) and about half the stopping power of BGO (1/tau=2.3 cm at 511 keV). Unlike CsF and NaI(T1), BaF.sub.2 is not hygroscopic.
The fast component of BaF.sub.2 emits light in the ultraviolet region of the spectrum. Glass photomultiplier tubes are not transparent to ultraviolet light, so a quartz photomultiplier tube must be used instead to detect the fast component of BaF.sub.2. Since quartz photomultiplier tubes are substantially more expensive than glass, one would prefer to avoid using BaF.sub.2, if possible, in favor of using a scintillator that can be detected by a glass photomultiplier tube. The fast component gives BaF.sub.2 very good timing resolution, but the slow component limits its high rate capabilities. In other words, it takes longer when using BaF.sub.2 to get ready for the next event.
Cerium fluoride is a relatively new scintillator material which has an excellent balance of properties. In comparison to NaI(T1), it has a much better stopping power (1/tau=1.9 cm instead of 3.0 cm), and its decay constant is almost ten times faster. The light yield is much lower (4 as compared to 100), but it is still sufficiently large to be useful.
Of the best known scintillator materials, BGO has the best stopping power, NaI(T1) has the best light yield, and BaF.sub.2 has the best timing resolution. However, as noted above, each of these materials have significant shortcomings which hinder their performance as scintillators for PET: BGO has a very long decay constant, NaI(T1) has a very long decay constant and is hygroscopic, and BaF.sub.2 has a long decay constant and requires expensive photomultiplier tubes. CeF.sub.3 provides good stopping power and good timing resolution, but since it was discovered little more than one year ago, it has yet to be incorporated in many radiation detectors and so its full potential is not yet known.